The present invention relates generally to a thermal energy transfer device for use within an x-ray generating device and, more specifically, to a convection cooled anode target for use within an x-ray tube.
Typically, an x-ray generating device, referred to as an x-ray tube, includes opposed electrodes enclosed within a cylindrical vacuum vessel. The vacuum vessel is commonly fabricated from glass or metal, such as stainless steel, copper, or a copper alloy. The electrodes include a cathode assembly positioned at some distance from the target track of a rotating, disc-shaped anode assembly. Alternatively, such as in industrial applications, the anode assembly may be stationary. The target track, or impact zone, of the anode is generally fabricated from a refractory metal with a high atomic number, such as tungsten or a tungsten alloy. Further, to accelerate electrons used to generate x-rays, a voltage difference of about 60 kV to about 140 kV is commonly maintained between the cathode and anode assemblies. The hot cathode filament emits thermal electrons that are accelerated across the potential difference, impacting the target zone of the anode assembly at high velocity. A small fraction of the kinetic energy of the electrons is converted to high-energy electromagnetic radiation, or x-rays, while the balance is contained in back-scattered electrons or converted to heat. The x-rays are emitted in all directions, emanating from a focal spot, and may be directed out of the vacuum vessel along a focal alignment path. In an x-ray tube having a metal vacuum vessel, for example, an x-ray transmissive window is fabricated into the vacuum vessel to allow an x-ray beam to exit at a desired location. After exiting the vacuum vessel, the x-rays are directed along the focal alignment path to penetrate an object, such as a human anatomical part for medical examination and diagnostic purposes. The x-rays transmitted through the object are intercepted by a detector or film, and an image of the internal anatomy of the object is formed. Likewise, industrial x-ray tubes may be used, for example, to inspect metal parts for cracks or to inspect the contents of luggage at an airport.
Since the production of x-rays in a medical diagnostic x-ray tube is by its very nature an inefficient process, the components in the x-ray tube operate at elevated temperatures. For example, the temperature of the anode""s focal spot may run as high as about 2,700 degrees C., while the temperature in other parts of the anode may run as high as about 1,800 degrees C. The thermal energy generated during tube operation is typically transferred from the anode, and other components, to the vacuum vessel. The vacuum vessel, in turn, is generally enclosed in a casing filled with a circulating cooling fluid, such as dielectric oil or air, that removes the thermal energy from the x-ray tube. The casing also supports and protects the x-ray tube and provides a structure for mounting the tube. Additionally, the casing is commonly lined with lead to shield stray radiation.
As discussed above, the primary electron beam generated by the cathode of an x-ray tube deposits a large heat load in the anode target. In fact, the target glows red-hot in operation. Typically, less than 1% of the primary electron beam energy is converted into x-rays, the balance being converted to thermal energy. This thermal energy from the hot target is conducted and radiated to other components within the vacuum vessel. The fluid circulating around the exterior of the vacuum vessel transfers some of this thermal energy out of the system. However, the high temperatures caused by this thermal energy subject the x-ray tube components to high thermal stresses that are problematic in the operation and reliability of the x-ray tube. This is true for a number of reasons. First, the exposure of components in the x-ray tube to cyclic high temperatures may decrease the life and reliability of the components. In particular, the anode assembly is subject to thermal growth and target burst. The anode assembly also typically includes a shaft that is rotatably supported by a bearing assembly. The bearing assembly is very sensitive to high heat loads. Overheating of the bearing assembly may lead to increased friction, increased noise, and to the ultimate failure of the bearing assembly. Due to the high temperatures present, the balls of the bearing assembly are typically coated with a solid lubricant. A preferred lubricant is lead, however, lead has a low melting point and is typically not used in a bearing assembly exposed to operating temperatures above about 330 degrees C. Because of this temperature limit, an x-ray tube with a bearing assembly including a lead lubricant is limited to shorter, less powerful x-ray exposures. Above about 450 degrees C., silver is generally the lubricant of choice, allowing for longer, more powerful x-ray exposures. Silver, however, increases the noise generated by the bearing assembly.
The high temperatures encountered within an x-ray tube also reduce the scanning performance or throughput of the tube, which is a function of the maximum operating temperature, and specifically the anode target and bearing temperatures, of the tube. As discussed above, the maximum operating temperature of an x-ray tube is a function of the power and length of x-ray exposure, as well as the time between x-ray exposures. Typically, an x-ray tube is designed to operate at a certain maximum temperature, corresponding to a certain heat capacity and a certain heat dissipation capability for the components within the tube. These limits are generally established with current x-ray routines in mind. However, new routines are continually being developed, routines that may push the limits of existing x-ray tube capabilities. Techniques utilizing higher power, longer x-ray exposures, and increased patient throughput are in demand to provide better images and greater patient care. This is especially true with respect to computed tomography (CT) systems. Thus, there is a need to remove as much heat as possible from existing x-ray tubes, as quickly as possible, in order to increase x-ray exposure power and duration before reaching tube operational limits.
The prior art has primarily relied upon removing thermal energy from the x-ray tube target by radiating heat from the target to the vacuum vessel wall and then transferring this heat to the cooling fluid circulating around the vacuum vessel. It has also relied upon increasing the diameter and mass of the anode target in order to increase the heat storage capability and radiating surface area of the target. These approaches have been marginally effective, however they are limited. The cooling fluid methods, for example, are not adequate when the anode end of the x-ray tube cannot be sufficiently exposed to the circulating fluid. This is a common problem in x-ray tubes having mounting and adjustment mechanisms. Other cooling fluid methods have sought to aid in the removal of heat from the x-ray tube by circulating fluid through multiple hollow chambers in the shaft of the anode assembly. These methods too are typically limited to hard-mounted x-ray tubes. Likewise, the target modification methods are generally not adequate as the potential diameter of the anode target is ultimately limited by space constraints on the scanning system, especially when enhanced x-ray system angulation capability is desired. Further, a finite amount of time is required for heat to be conducted from the target track, where the electron beam actually hits the anode target, to other regions of the target. In fact, thermal energy may not even reach the back of the target until a given scan has ended. Thus, adding extra mass to the back of the target provides little thermal performance benefit.
Therefore, what is needed are devices providing enhanced anode target heat dissipation, thus enabling lower target track and bulk temperatures, enabling higher peak power for a given x-ray tube rotor speed, reducing the risk of target burst, and allowing longer and more powerful x-ray scans. What is also needed are devices providing smaller targets with lower target mass for a given power rating, for example, decreasing the bearing load on CT tubes, enabling higher CT system gantry speeds, and allowing better x-ray system angulation.
The present invention overcomes the aforementioned problems and permits greater x-ray tube throughput by providing a cooler anode target. The present invention also reduces thermal growth of the anode target, improving image quality and allowing for the simplification of CT system design. Further, the present invention increases the life of x-ray tube components.
In one embodiment, an anode assembly for use within an x-ray generating device includes a target frame having an inner surface and an outer surface; a heat exchanger having an inner surface and an outer surface, at least a portion of the outer surface of the heat exchanger positioned adjacent to at least a portion of the inner surface of the target frame; and a thermal coupling medium disposed between the inner surface of the target frame and the outer surface of the heat exchanger, the thermal coupling medium thermally coupling the target frame with the heat exchanger.
In another embodiment, a thermal energy transfer device for use within a target of an anode assembly of an x-ray generating device includes a heat exchanger having an inner surface and an outer surface, at least a portion of the outer surface of the heat exchanger positioned adjacent to at least a portion of an inner surface of a target frame of the anode target; and a thermal coupling medium disposed between the inner surface of the target frame and the outer surface of the heat exchanger, the thermal coupling medium thermally coupling the target frame with the heat exchanger while permitting relative motion between the target frame and the heat exchanger.
In a further embodiment, an x-ray generating device that generates x-rays and residual energy in the form of heat includes, a vacuum vessel having an inner surface forming a vacuum chamber; an anode assembly disposed with the vacuum chamber, the anode assembly including a target having a target frame with an inner surface and an outer surface; a cathode assembly disposed within the vacuum chamber at a distance from the anode assembly, the cathode assembly configured to emit electrons that strike the target, producing x-rays and residual energy; a heat exchanger having an inner surface and an outer surface, at least a portion of the outer surface of the heat exchanger positioned adjacent to and in a spaced apart relationship with at least a portion of the inner surface of the target frame; and a thermal coupling medium disposed between the inner surface of the target frame and the outer surface of the heat exchanger, the thermal coupling medium thermally coupling the target frame with the heat exchanger while permitting relative motion between the target frame and the heat exchanger.